Diabetes Mellitus is a disease of major global importance that has increased in frequency at almost epidemic rates. The worldwide prevalence in 2006 is 170 million people and predicted to at least double over the next 10-15 years. Diabetes is characterized by a chronically raised blood glucose concentration (hyperglycemia), due to a relative or absolute lack of the pancreatic hormone, insulin. Within the healthy pancreas, beta cells, located in the islets of Langerhans, continuously produce and secrete insulin according to the blood glucose levels and thus maintain near constant glucose levels in the body.
Much of the burden of the disease to the patient and to health care resources is due to the long-term tissue complications, which affect both the small blood vessels (microangiopathy, causing eye, kidney and nerve damage) and the large blood vessels (causing accelerated atherosclerosis, with increased rates of coronary heart disease, peripheral vascular disease and stroke). The Diabetes Control and Complications Trial (DCCT) demonstrated that development and progression of the chronic complications of diabetes are greatly related to the degree of altered glycemia as quantified by determinations of glycohemoglobin. [DCCT Trial, N. Engl. J. Med. 1993; 329: 977-986, UKPDS Trial, Lancet 1998; 352: 837-853. BMJ 1998; 317, (7160): 703-13 and the EDIC Trial, N. Engl. J. Med. 2005; 353, (25): 2643-53]. Thus, maintaining euglycemia by frequent glucose measurements and adjustment of insulin delivery accordingly is of utmost importance.
In theory, returning blood glucose levels to normal by hormone replacement therapy using insulin injections and/or other treatments in diabetes should prevent complications, but, frustratingly, near-normal blood glucose concentrations are very difficult to achieve and maintain in many patients, particularly those with type 1 diabetes. In these patients, blood glucose concentration can swing between very high (hyperglycemia) and very low (hypoglycemia) levels in an unpredictable manner. Thus, tight glycemic control is required. This control can be achieved by providing a means capable of substituting the two functions of the normal pancreas—glucose monitoring and insulin delivery. Furthermore, a closed loop system provided with a feedback mechanism linking between both functions (often referred to as an “artificial pancreas”) could theoretically maintain near normal blood glucose levels.
Continuous subcutaneous insulin infusion (CSII) via pumps provides a closer approximation of normal plasma insulin profiles and increased flexibility regarding timing of meals and snacks compared to conventional insulin injection regimens.
In addition, insulin pump therapy in Diabetes Mellitus is associated with improved metabolic control and reduced risk of severe hypoglycemia compared to multiple daily injections of insulin.
Several ambulatory insulin infusion devices are currently available on the market. First generation pumps fitted with disposable syringe-type reservoirs and tubes are described by in U.S. Pat. No. 3,631,847 to Hobbs, U.S. Pat. No. 3,771,694 to Kaminski, U.S. Pat. No. 4,657,486 to Julius, and U.S. Pat. No. 4,544,369 to Skakoon. The main drawbacks of these devices are their large size and the weight, caused by the spatial configuration and the relatively large driving mechanism of the syringe and the piston. The relatively bulky device is carried in a patient's pocket or attached to the belt or some other article of clothing. Consequently, the fluid delivery tube is typically quite long, usually longer than 60 cm, to permit needle insertion in remote sites of the body. These uncomfortable, bulky devices with a long tube are rejected by the majority of diabetic insulin users, because they tend to disturb regular activities, such as for example sleeping and swimming. Furthermore, the effect of the image projected on a teenagers' body is unacceptable. In addition, the presence of the delivery tube excludes some otherwise potentially preferable remote insertion sites, like the buttocks and the extremities.
Pumps can be provided with a housing having a bottom surface adapted for contact with the patient's skin, a reservoir disposed within the housing, and an injection needle adapted for communication with the reservoir. These skin adhered pumps should be disposed of every 2-3 days like current pump infusion sets. This paradigm was described by Burton in U.S. Pat. No. 5,957,895, Connelly, in U.S. Pat. No. 6,589,229, and by Flaherty in U.S. Pat. No. 6,740,059. Other configuration of the skin adhered pumps are disclosed in U.S. Pat. Nos. 6,723,072 and 6,485,461. The pump can include a single piece that adheres to the patient skin for the entire usage duration. The needle emerges from the bottom surface of the pump and is fixed to the device housing. These so-called “second-generation” skin adhered devices tend to be expensive, bulky and heavy.
Current diabetic patients generally measure their own blood glucose level discontinuously, on the order of perhaps several times during the day. Blood glucose sampling typically includes obtaining finger-prick capillary samples and applying the blood to a reagent strip for analysis done in a portable meter. The discomfort involved with these methods often leads to poor patient compliance. Testing cannot be performed while sleeping and while the subject is occupied. In addition, the results do not give information regarding trends in glucose levels, but rather provide only discrete readings, taken at typically large time intervals between consecutive measurements. Therefore it would be desirable to carry out the glucose monitoring substantially continuously by performing discrete measurements, at a very high rate. Continuous monitoring can be done by invasive, minimally-invasive, or non-invasive means.
Minimally-invasive glucose monitors can measure glucose levels in the interstitial fluid (ISF) present within the subcutaneous tissue. The strong correlation between blood and ISF glucose levels has been shown to facilitate accurate glucose measurements (Diabetologia 1992; 35, (12): 1177-1180).
The GlucoWatch G2® Biographer (available from Cygnus, Inc., Redwood City, Calif.) is one commercially available minimally-invasive glucose monitor whose function is detailed in U.S. Pat. No. 6,391,432. A small current passing between two skin-surface electrodes draws ions and (by electro-endosmosis) glucose-containing interstitial fluid to the skin-surface and into hydrogel pads provided with a glucose oxidase (GOX) biosensor (JAMA 1999; 282: 1839-1844). Readings are taken every 10 min, with a single capillary blood calibration. Disadvantages of the GlucoWatch® are associated with occasional sensor values differing markedly from blood values; with skin rashes and irritation in those locations which are immediately underneath the device, appearing in many users; with a long warm up time of 3 hours; and with skips in measurements due to sweating.
Another commercially available minimally-invasive monitor is the Guardian® RT Continuous Glucose Monitoring System (available from Medtronic MiniMed Inc., Northridge, Calif.). This device is a GOX-based sensor, which is described in U.S. Pat. No. 6,892,085. The sensor consists of a subcutaneously implanted, needle-type, amperometric enzyme electrode, coupled with a portable logger (Diab. Tech. Ther. 2000; 2: Supp. 1, 13-18). The Guardian® RT system displays updated glucose readings every five minutes, together with hypo- and hyperglycemic alarms. The sensor is based on the technology of GOX immobilized at a positively charged base electrode, with electrochemical detection of hydrogen peroxide production. This enzymatic reaction, when carried out in-vivo, can encounter stoichiometric hurdles that can compromise its accuracy. The device is large and bulky and requires inconvenient tubing.
Closed loop infusion systems, such as the system described in U.S. Pat. No. 6,558,351, can include a sensor system and a delivery system. The systems can be coupled via a controller that uses the inputs of the sensor system to generate commands to the delivery system. The main shortcoming of the described closed loop system is that two separate devices, comprising separate tubing and separate cannulae should be applied to the body of the user.
Measurement of the glucose concentration in a sample can be performed using one or more methods. The most common methods today are electrochemical measurement techniques and optical measurement techniques. The detection principle of enzyme-based sensors is based on monitoring of the enzyme-catalyzed oxidation of glucose. These include glucose sensors using amperometric or potentiometric operating principles.
The enzymatic reaction that occurs in the majority of these sensors is catalyzed by glucose oxidase (GOX). In this reaction, oxygen and glucose yield gluconic acid and hydrogen peroxide. In this reaction, in which glucose is oxidized to gluconic acid, glucose oxidase acts temporarily as an electron acceptor, which means that it is first reduced to an inactive state and is subsequently reactivated by the reduction of oxygen to hydrogen peroxide. The analyte concentration is transduced into a detectable signal, generally by using amperometric methods.
Ex vivo amperometric glucose sensors are often mediator-based. A mediator-based glucose sensor uses an artificial electron acceptor, or mediator, to replace the natural acceptor, oxygen, in the oxidation of glucose by GOX and thus is not oxygen dependent. The oxidation of the reduced mediator occurs at a low potential, thus reducing the sensitivity of the sensor to interfering substances.
Many in vivo devices are mediatorless due to possible leaching and toxicity of the mediator. Illustrative examples of mediatorless devices are described in U.S. Publication No. 2005/0272989 assigned to MiniMed, and U.S. Pat. No. 6,975,893 assigned to Abbott. These devices generally rely on oxygen as a physiological electron acceptor. Arterial blood has a glucose-to-oxygen ratio of approximately 10 to 1, while venous blood has a glucose-to-oxygen ratio of about 100 to 1. Thus, in vivo devices often use membranes to tailor the flux of glucose and oxygen to the enzymatic coating on the electrode. Different layers alter the diffusion of one or more analytes into the area that comprises the catalytic enzyme or enzymes. The stoichiometric scarcity of oxygen in vivo can present an obstacle to the effectiveness of these devices. In addition, these devices are subjected to errors due to fluctuations in the concentration of dissolved oxygen.
Amperometric measurement of hydrogen peroxide requires application of a potential at which additional electroactive species, such as for example ascorbic and uric acids or acetaminophen are present. These and other oxidizable constituents of biological fluids can compromise the selectivity and hence the overall accuracy of the glucose concentration measurement. Hydrogen peroxide can have toxic effects that may compromise the biocompatibility of the sensor. This poses a problem mainly when the hydrogen peroxide is not consumed for the transduction (that is, when the biosensor is not based on hydrogen peroxide). Application of catalase may resolve this setback. Hydrogen peroxide also tends to deactivate the GOX molecules, limiting the time available for application of the sensor. Overloading the sensor with an excess of enzyme, more than what is required to catalyze the incoming glucose, may be helpful in overcoming this problem. Co-immobilization of catalase may be beneficial. However, this solution is more appropriate for glucose sensors based on the detection of O2 that do not depend on measuring H2O2. Furthermore, catalase is in turn inactivated by hydrogen peroxide (Diab. Tech. & Ther. Vol. 2, No. 3, 2000, pp. 367-376). Additionally, the size of such probes, including the sensing unit with its various layers, is relatively large, affecting the ease and comfort of the probe insertion into the user's body. Miniaturizing the sensing technology within the probe, which requires high levels of enzyme loading, while keeping high measurement sensitivity, remains a challenge.
Microdialysis is an additional commercially available minimally-invasive technology (Diab. Care 2002; 25: 347-352) for glucose monitoring as detailed in U.S. Pat. No. 6,091,976 to Pfeiffer (assigned to Roche Diagnostics, Basel, Switzerland) and used in the GlucoDay® S (available from Menarini Diagnostics, Florence, Italy). A fine, semi-permeable hollow dialysis fiber is implanted in the subcutaneous tissue and perfused with isotonic fluid. Glucose diffuses across the semi-permeable fiber and is pumped outside the body via the microdialysis mechanism for measurement by a glucose oxidase-based electrochemical sensor. Initial reports (Diab. Care 2002; 25: 347-352) show good agreement between sensor and blood glucose readings, and good stability with a one-point calibration over one day. Higher accuracies were found when using the microdialysis-based sensor, compared to the needle-type sensor (Diab. Care 2005; 28, (12): 2871-6).
Disadvantages of the microdialysis-based glucose sensors stem primarily from the constant perfusion of solution through the microdialysis probe. This operational method requires the presence of a dedicated pump and reservoir, leading to large and bulky devices, and also necessitates high energy consumption. Furthermore, the relatively large size of the microdialysis catheter often causes a wound and subsequent local tissue reactions following its insertion into the subcutaneous tissue. Finally, the microdialysis process generates long measurement lag times, due to the essential slow perfusion rates and long tubing.
Optical glucose measurement techniques can be attractive for several reasons: they utilize non-ionizing electromagnetic radiation to interrogate the sample, they do not generally require consumable reagents, and they are fast. Also, these techniques are generally non-destructive and reagentless, thereby reducing the risk of unsafe reactions and their byproducts. Although optical approaches for glucose sensing are attractive, they can nevertheless be plagued by a lack of sensitivity and/or specificity since variations in optical measurements depend on variations of many factors in addition to glucose concentration. Isolating those changes which are due to glucose alone and using them to predict glucose concentration is a significant challenge in itself (J. Biomedical Optics 5 (1), 5-16 Jan. 2000). Furthermore, non-invasive optical glucose monitors, which involve sensing of glucose concentration levels through the skin, involve very low signal-to-noise ratio, scattering and interferences by bodily fluids and by the skin itself, causing non-invasive optical sensors to lack specificity and repeatability.